This invention relates to the field of tomography. More specifically, the present invention relates to a method of measuring and correcting the attenuation associated with detecting coincidences using a collimated source and a dedicated detector for improved measurement sensitivity.
Positron Emission Tomography (PET) has gained significant popularity in nuclear medicine because of the ability to non-invasively study physiological processes within the body. Applications employing the PET technology for its sensitivity and accuracy include those in the fields of oncology, cardiology and neurology.
Using compounds such as 11C-labeled glucose, 18F-labeled glucose, 13N-labeled ammonia and 15O-labeled water, PET can be used to study such physiological phenomena as blood flow, tissue viability, and in vivo brain neuron activity. Positrons emitted by these neutron deficient compounds interact with free electrons in the body area of interest, resulting in the annihilation of the positron. This annihilation yields the simultaneous emission of a pair of photons (gamma rays) approximately 180xc2x0 (angular) apart. A compound having the desired physiological effect is administered to the patient, and the radiation resulting from annihilation is detected by a PET tomograph. After acquiring these annihilation xe2x80x9cevent pairsxe2x80x9d for a period of time, the isotope distribution in a cross section of the body can be reconstructed.
PET data acquisition occurs by detection of both photons emitted from the annihilation of the positron in a coincidence scheme. Due to the approximate 180xc2x0 angle of departure from the annihilation site, the location of the two detectors registering the xe2x80x9ceventxe2x80x9d define a chord passing through the location of the annihilation. By histogramming these lines of response (the chords), a xe2x80x9csinogramxe2x80x9d is produced that may be used by a process of back-projection to produce a three dimensional image of the activity. Detection of these lines of activity is performed by a coincidence detection scheme. A valid event line is registered if both photons of an annihilation are detected within a coincidence window of time. Coincidence detection methods ensure (disregarding other second-order effects) that an event line is histogrammed only if both photons originate from the same positron annihilation.
In the traditional (2-D) acquisition of a modern PET tomograph, a collimator (usually tungsten) known as a septa is placed between the object within the field-of-view and the discrete axial rings of detectors. This septa limits the axial angle at which a gamma ray can impinge on a detector, typically limiting the number of axial rings of detectors that a given detector in a specific ring can form a coincidence with to a few rings toward the front of the tomograph from the given detector""s ring, the same ring that the detector is within, and a few rings toward the rear of the tomograph from the given detector""s ring.
The current state of the art nuclear medicine camera is capable of operating in either PET or Single Photon Emission Computed Tomography, (xe2x80x9cSPECTxe2x80x9d), modes. Fundamentally, in both modes, gamma photons emitted from within a patient are detected. However, a significant portion of these emitted photons are obstructed from reaching the detectors by colliding with atoms. When this occurs, one significant possibility is a course alteration away from the detector that may result in a missed detection. The degree of attenuation depends upon the amount and density of matter between the emitting source and the detector, and will vary from subject to subject. The more attenuation present, the less probable will be the accurate detection of a gamma photon. Unless the amount of attenuation is known, the detected activity within a defined energy window underestimates the true activity. This results in poorer contrast and attenuation artifacts in the reconstructed images. Conditions such as these reduce the confidence one may have in extracting information for diagnosis. However, attenuation methods well suited for SPECT are poorly suited for PET.
To compensate for this phenomenon, it is now common to incorporate an apparatus to transmit gamma photons of a known flux density through a patient so that the patient induced attenuation can be measured. Attenuation was first measured in PET by using a ring of positron emitting isotope surrounding the object to be measured. In this technique, the ratio between a transmission scan and a blank scan form the attenuation. The blank is measured by simply measuring the rate that gamma rays from positrons are detected by the detection system when no attenuating media is present. In the original scanners as described above as having septa, the septa are provided for collimating the gamma rays in an axial direction, but the rings allow for no transaxial collimation. The lack of collimation allow the acceptance of scattered events into the transmission measurement, resulting in an underestimate of the attenuation. To improve the transmission measurement, systems use rotating rod sources. These sources are disposed in parallel fashion to the axis of the scanner and are collimated in the axial direction by the septa. In the transaxial direction, the collimation may be provided electronically since the position of the source is known. However, the activity in the rod must be the same as that activity in the earlier ring source to provide the same count rate. With modern block detectors, the dead-time of the near block limits the activity in the rod.
A more recent advancement in PET acquisition is 3-D, in which the septa are removed, which allows a given detector to be in coincidence with detectors from all other detector rings. With the advent of three-dimensional reconstruction techniques, greater sensitivity to emission counts is possible if the septa are removed. As the septa represent a significant cost, there is also an economic incentive to exclude them from the system. However, with the absence of septa, the problems of both detector dead-time and scatter are magnified.
Since the position of a source with respect to the detector system can be known, there is no need to detect coincidences, thereby allowing the use of a source that emits single gamma rays. Only one detectorxe2x80x94the detector on the far side of the systemxe2x80x94is needed to make the transmission or blank measurements. Without the counting losses due to the dead-time of the near detector, the activity of the source may be increased resulting in an increase in count-rate and thus a better quality measurement. However, without axial collimation, the scatter included in the transmission scan causes an underestimate of the attenuation measurement. To decrease the possibility of scatter, the gamma rays from the source can be collimated with lead or tungsten to form a beam that illuminates only a narrow plane of detectors. Other gamma rays that would only contribute to background are eliminated. Since the directionality of single gamma rays cannot be determined, only a single point of activity illuminating a detector bank can be used. This requires increased levels of activity to meet the count-rate needed for an adequate quality measurement. Also, the scanning protocol is more efficient if the transmission measurement is performed after the patient has been injected with radioactivity. Even though a different isotope such as 137Cs which emits gamma rays with an energy of 662 keV can be used for the transmission scan, there is a significant difficulty in distinguishing the transmission events from the emission events.
(SPECT) is similar to PET. However, in SPECT, only a single photon from a nuclear decay within the patient is detected. Also, the line of response traveled by the photon is determined exclusively by detector collimation in SPECT, as opposed to the coincident detection of two collinear photons as in PET.
There are a number of correction methods in the art. However, these methods utilize SPECT based approaches for transmission based attenuation correction that is not well suited to PET. Typical of the art are the following:
Also typical of the art are the following:
Tan P., Bailey D. L., Meikle S. R., Eberl S. Fulton R. R., and Hutton B. F., A Scanning Line Source for Simultaneous Emission and Transmission Measurements in SPECT, Journal of Nuclear Medicine, col. 34, No. 10, October 1993.
Lange K., Carson R., EM Reconstruction Algorithms for Emission and Transmission Tomography, Journal of Computer Assisted Tomography, Vol and No. 2, 1984.
Bailey D. L., Hutton B. F., Walker P. J., Improved SPECT Using Simultaneous Emission and Transmission Tomography, Journal of Nuclear Medicine, Vol. 28, No. 5, May 1987.
Lange K., Bahn M., Little R., A Theoretical Study of Some Maximum Likelihood Algorithms for Emission and Transmission Tomography, IEEE Log Number 8714498, 1987.
Gullberg G. T., Huesman R. H., Malko R. A., Plec N. J., Budinger T. F., An Attenuated Proiector-Backprojector for Iterative SPECT Reconstruction, Physics in Medicine and Biology, Vol. 30, No. 8, 799-816, 1985.
Chang L. T., A Method for Attenuation Correction in Radionuclide Computed Tomography, IEEE Transactions for Nuclear Science, Vol. NS-25, No. 1, February 1978.
Hudson H. M., Larkin R. S., Accelerated Image Reconstruction Using Ordered Subsets of Projection Data, IEEE Transactions on Medical Imaging, Vol. 13, No. 4, December 1994.
Hollinger D. F., Loncaric S., Yu D. C., Ali A., Chang W., Using Fast Sequential Asymetric Fanbeam Transmission CT for Attenuation Correction of Cardiac SPECT Imaging, Journal of Nuclear Medicine, 1998; 39:1335-1344.
Kak A. C., Slaney M., Principles of Computerized Tomographic Imaging, IEEE Press 1987 ISBN 0-7803-0447-0.
Sorenson J. A., Phelps M. E., Physics in Nuclear Medicine, Second Edition, Grune and Stratton, Inc. Harcourt, Brace Jovanovich, 1987, ISBN 0-8089-1804-4.
The details of carrying out a PET study are given in numerous publications. Typically, the following references provide a background for PET. These are incorporated herein by reference for any of their teachings.
1. M. E. Phelps, et al.: xe2x80x9cPositron Emission Tomography and Audiographyxe2x80x9d, Raven Press, 1986;
2. R. D. Evans: xe2x80x9cThe Atomic Nucleusxe2x80x9d, Kreiger, 1955;
3. J. C. Moyers: xe2x80x9cA High Performance Detector Electronics System for Positron Emission Tomographyxe2x80x9d, Masters Thesis, University of Tennessee, Knoxville, Tenn., 1990;
4. U.S. Pat. No. 4,743,764 issued to M. E. Casey, et al, on May 10, 1988;
5. R. A. DeKemp, et al.: xe2x80x9cAttenuation Correction in PET Using Single Photon Transmission Measurementxe2x80x9d, Med. Phys., vol. 21, 771-8, 1994;
6. S. R. Cherry, et al.: xe2x80x9c3-D PET Using a Conventional Multislice Tomograph Without Septaxe2x80x9d, JI. C. A. T., 15(4) 655-668.
7. J. S. Karp, et al.: xe2x80x9cSingles Transmission in Volume-Imaging PET With a 137Cs Sourcexe2x80x9d, Phys. Med. Biol. Vol. 40, 929-944 (1995).
8. S. K. Yu, et al.: xe2x80x9cSingle-Photon Transmission Measurements in Positron Tomography Using 137Csxe2x80x9d, Phys. Med. Biol. Vol. 40, 1255-1266 (1995).
9. G. F. Knoll: Radiation Detection and Measurement, John Wiley and Sons (1989).
10. S. R. Cherry, et al.: xe2x80x9cOptical Fiber Readout of Scintillator Arrays using a Multi-Channel PMT: A High Resolution PET Detector for Animal Imagingxe2x80x9d,. IEEE Transactions on Nuclear Science, Vol. 43, No. 3, 1932-1937 (June, 1996).
11. J. A. McIntyre, et al.: xe2x80x9cConstruction of a Positron Emission Tomograph with 2.4 mm Detectorsxe2x80x9d, IEEE Transactions on Nuclear Science, Vol. 33, No. 1,425-427 (February, 1986).
Both PET and SPECT systems are also well known to persons skilled in the art.
In order to achieve maximal quantitative measurement accuracy in tomography applications, an attenuation correction must be applied to the collected emission data. In a PET system, for example, this attenuation is dependent on both the total distance the two gamma rays must travel before striking the detector, and the density of the attenuating media in the path of travel. Depending on the location of the line of response within. the patient""s body, large variations in attenuating media cross section and density have to be traversed. If not corrected for, this attenuation causes unwanted spatial variations in the images that degrade the desired accuracy. As an example, for a cardiac study the attenuation is highest in the line of responses (LORs) passing through the width of the torso and arms, and attenuation is lowest in the LORs passing through from the front to the back of the chest.
Typically, the attenuation correction data in PET systems is produced by either: shape fitting and linear calculations using known attenuation constants, these being applicable to symmetric well-defined shapes such as the head and torso below the thorax (calculated attenuation); or through the measurement of the annihilation photon path""s attenuation using a separate transmission scan (measured attenuation). The use of calculated attenuation correction, which introduces no statistical noise into the emission data, can be automated for simple geometries such as the head, and is the most prominent method used for brain studies. However, complexities in the attenuation media geometry within the chest have prevented the application of calculated attenuation from being practical for studies within this region of the body. Accordingly, transmission scanning has been utilized.
The total attenuation of a beam along a LOR through an object is equal to the attenuation that occurs for the two photons from an annihilation. Thus, the emission attenuation along the path can be measured by placing a source of gamma rays on the LOR outside of the body and measuring attenuation through the body along this line. It has been the practice to accomplish this attenuation measurement by placing a cylindrical positron emitter xe2x80x9csheetxe2x80x9d within the PET tomograph""s field of view (FOV) but outside of the region (the object) to be measured. The ratio of an already acquired blank scan (no object in the FOV) to the acquired transmission scan is calculated. These data represent the desired measured attenuation factors, which may vary spatially. These data are then applied to the emission data after a transmission scan of the object to correct for the spatial variations in attenuation.
There are two types of transmitter source units conventionally utilized in PET transmission scan data collection, both of which form a xe2x80x9csheetxe2x80x9d of activity to surround the patient. One involves the placement of rings of activity aligned with detector rings around the inner face of the septa. The second type utilizes the rotation of one or more axially-oriented rods of activity in a circular path just inside the inner face of the septa.
The first of these two emitter systems (the ring source method) significantly reduces the sensitivity of the tomograph due to the close source-proximity dead time effects of the source activity on all of the detectors. Further, removal of this assembly is either performed manually by facility personnel or by a complex automated mechanical assembly. Large, cumbersome, out of the FOV shielding is required for storage of the automated source when not in use, adding to the depth of the tomograph tunnel and, thus increasing incidence of patient claustrophobia. The second type of emitter, using rotating source(s) suffers from the above-mentioned problems and also, due to the shielding requirements, reduces the patient tunnel diameter, further increasing patient claustrophobia symptoms.
Both of the above automated source transportation methods suffer from high mechanical component cost and from low sensitivity. Due to the dead-time-induced reduction in tomograph sensitivity, lengthy acquisitions are required in order to achieve usable low noise transmission scan data.
In order to reduce costs in scintillator detector applications, multiplexing techniques based on the use of fiber optics are advantageous. Those disclosures made by Cherry, et al. (Cherry), and McIntyre, et al. (McIntyre), teach the use of fiber optics connected between the imaging detectors and multichannel photomultipliers (PMT""s). Cherry discloses the use of a multi-channel PMT in association with an 8xc3x978 array of bismuth germanate (BGO) crystals. As discussed by Cherry, a charge division readout board is used to convert the 64 signals into four position sensitive signals which determine the crystal interaction. In the earlier McIntyre article, the authors disclose the use of fiber optics coupled between the detectors and a number of multi-channel PMT""s. Specifically, McIntyre teaches the use of 288 PMT""s in association with 8,192 detectors, for reducing the number of required PMT""s by a factor of about 28.4.
In the McIntyre embodiment, eight detector rings are each divided into four quadrants. Each ring is comprised of sixteen concentric rings. The respective quadrants for the eight detector rings are grouped together for a total of 256 detectors per quadrant group. Sixteen xe2x80x9ccoarsexe2x80x9d fiber sets connect sixteen PMT""s to the 256 detectors, with sixteen detectors in one ring quadrant connected to one PMT. Similarly, sixteen xe2x80x9cfinexe2x80x9d fiber sets connect sixteen PMT""s to the 256 detectors, with corresponding detectors in each ring quadrant of a quadrant group being connected to one PMT. One PMT is connected to each ring quadrant. Thus, a total of 32 PMT""s are required for determining the particular detector xe2x80x9c"THgr"xe2x80x9d address within a quadrant. Similarly, 32 PMT""s are required to determine the xe2x80x9crxe2x80x9d address, corresponding to which of the concentric rings in a particular ring the detector is disposed. Finally, eight PMT""s are required to determine which ring quadrant the detector is disposed. Thus, a total of 72 PMT""s are required for each quadrant for a total of 288 PMT""s in association with 8,192 detectors.
Therefore, it is an object of the present invention to provide a system for detecting coincident activity from a point source.
Another object of the present invention is to provide such a system which includes a detector dedicated to collecting attenuation data.
Yet another object of the present invention is to provide a system for detecting coincident activity while illuminating only a strip of the imaging detector in order to eliminate events not of interest in the attenuation measurement.
A further object of the present invention is to provide a collimated point source and dedicated detector whereby only a selected strip of the imaging detector is illuminated such that events unrelated to the attenuation are eliminated.
Still another object of the present invention is to provide an arrangement whereby gamma radiation detected by dedicated detectors is transmitted to a plurality of PMT""s such that an address of each gamma radiation detector is readily determined and such that the total required number of PMT""s is reduced relative to conventional devices.
A further object of the present invention is to provide a single method for acquiring patient attenuation measurements on a dual modality PET/SPECT camera.
A still further object of the present invention is to provide a transmission source capable of xe2x80x9cshining throughxe2x80x9d the patient and the collimator.
It is yet another object of the present invention to provide a method and a device in which transmission data emission data can be acquired either simultaneously, gated sequential in which emission data then transmission data are acquired at each gantry angle, or sequentially in which a transmission scan is performed either before or after a complete emission subject study is collected.
A further object of the present invention is to provide a method and device in which the placement of the transmission source will allow a continuous range of angles between the source and the opposed detector ranging from 90 to 180 degrees.
Other objects and advantages of the present invention will become more apparent upon review of the detailed description and associated drawings of the scintillator detector array for encoding the energy, position and time coordinates of gamma-ray interactions.
Other objects and advantages will be accomplished by the present invention which serves to detect activity from a, preferably, collimated point source. The present invention includes a detector dedicated to collecting attenuation data. The collimated point source and dedicated detector are positioned with respect to the tomography device such that only a selected strip of the imaging detector is illuminated such that events unrelated to the attenuation are eliminated.
The source of the present invention includes a collimator in which is disposed a point source. An opening is defined by the collimator for exposing a selected portion of the imaging detectors of the tomograph device. Positioned behind the point source, relative to the imaging detectors, is an attenuation detector dedicated to collecting attenuation data. Because the attenuation detector is dedicated to the attenuation measurement, the requirements for the attenuation detector are different from those for the imaging detector. For instance, it is not required that the attenuation detector be able to accurately determine the energy or spatial position of events within the detector, as is necessary for standard imaging detectors. It is therefore possible to design such an attenuation detector with much less dead time, and much higher count rate performance, than a standard imaging detector. The improved count rate performance of the attenuation detector enables significant reduction of statistical noise in the attenuation correction measurement. The attenuation detector and collimator are designed to illuminate only a strip of the imaging detector, and the corresponding aperture of the attenuation detector, thereby eliminating events not of interest in the attenuation measurement. This also reduces dead time of the system and improves the count rate performance for events of interest.
A source of the present invention is disposed opposite each bank of imaging detectors of a dual head camera. Each source contains four point sources arranged along the axial extent. The sources and the associated collimators are positioned to the side of each head at a slight angle relative to the respective head. The sources and detectors are fixed relative to the imaging heads. In order to obtain full coverage of the field of view (FOV) in the same manner as for an emission scan, the heads and sources are rotated about the center of the camera.
The present invention further provides an arrangement of fiber optics interconnected between a plurality of dedicated gamma radiation detectors and a lesser number of photomultiplier tubes. The gamma radiation detectors are each provided for dedicated detection of 511 keV gamma radiation from one of a plurality of point sources disposed in a collimator. The arrangement of fiber optics is designed such that the address of a particular gamma radiation detector is readily discernable while minimizing the number of PMT""s required to process data accumulated by the gamma radiation detectors.
An alternate embodiment is described that allows provide a single method for acquiring patient attenuation measurements on a dual modality PET/SPECT camera. As recognized by those skilled in the art, dual purpose PET/SPECT cameras typically consist of two or more rectangular detectors. As described above, a collimated point source and dedicated detector are positioned with respect to the tomography device such that only a selected strip of the imaging detector is illuminated such that events unrelated to the attenuation are eliminated. Since state of the art SPECT requires a collimator in front of the detector, due to the current state of the art in gamma photon detection, the offset transmission source will likely not be in the direct FOV of the collimator. This results in attenuation to various degrees of the transmission beam (flux) due to the collimator. To overcome this problem, in the alternate embodiment, the point source is selected to have a sufficiently high energy and flux density to xe2x80x9cshine throughxe2x80x9d the subject and additionally, the collimator.